Chamber for encapsulating secreting cells

ABSTRACT

The invention relates to an encapsulating chamber for secreting cells, comprising a closed shell made of a semi-permeable membrane, said membrane comprising at least one layer of porous biocompatible polymer, and one layer of non-woven biocompatible polymer.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a U.S. National Stage Application pursuant to 35U.S.C. § 371 of International Patent Application PCT/EP2014/076955,filed on Dec. 9, 2014, and published as WO 2015/086550 on Jun. 18, 2015,which claims priority to French Patent Application 13/62342 filed onDec. 10, 2013, all of which are incorporated herein by reference intheir entireties for all purposes.

The invention relates to the field of bioartificial organs which areimplantable and in particular which are in the form of chambers forencapsulating cells secreting a substance of interest. The membraneswhich enable such encapsulating chambers and bioartificial organs to bemanufactured are also subjects of the invention.

The treatment of pathological conditions requiring a continuous supply,to the body, of substances of therapeutic interest has made necessarythe development of devices which can be implanted in a patient and arecapable of releasing these substances efficiently and sometimes for longperiods of time.

To satisfy this need, bioartificial organs which contain cells producingone or more substances of therapeutic interest have been developed. Thecells contained in a bioartificial organ are confined in internalspaces, or encapsulating chambers, delimited by at least onesemi-permeable membrane. Such a membrane is termed “semi-permeable” whenit allows the diffusion of the substances of therapeutic interest out ofthe encapsulating chamber to the target cells in the patient's body,while at the same time being impermeable to the antibodies and the cellsof the patient's immune system, thus preventing them from directlyattaching the cells producing the substance(s) of therapeutic interest.

A bioartificial organ is understood to be a device, in particularintended to be implanted in a patient, comprising at least oneencapsulating chamber consisting of at least one semi-permeablemembrane; said encapsulating chamber is intended to contain cells whichsecrete one or more substance(s) of therapeutic interest.

These substances of therapeutic interest are any substance intended tohave a beneficial effect in the patient. These may therefore be aneurotransmitter, a hormone, a growth factor, a coagulation factor or acytokine. In particular, these substances may be, without any limitingnature, insulin, glucagon, growth hormone, coagulation factor IX,coagulation cofactor VIII or calcitonin.

Examples of devices (bioartificial organs, semi-permeable membranes,encapsulating chambers) are known in the prior art.

Mention may thus be made of WO 02/060409 which describes a membraneconsisting of a porous polycarbonate biocompatible film which issurface-modified by generation of polar sites and covered with a layerof at least one hydrophilic polymer, and the use thereof formanufacturing bioartificial organs.

WO 2012/017337 and FR 2960783 describe a functionalized semi-permeablemembrane composed of a porous biocompatible support pretreated so as toincrease the surface energy thereof and comprising at least two layers,each comprising a hydrophilic polymer and at least one biologicallyactive molecule, and also the use thereof in particular formanufacturing a bioartificial organ and an encapsulation chamber.

The membrane disclosed in these documents doesn't present the two layers(porous biocompatible polymer and non-woven polymer) disclosed herein.This is clear in view of FIG. 2 of FR 2960783 which shows that thehydrophilic layers (3) have been deposited onto a unique layer of porousbiocompatible polymer (2). It is also to be noted that such hydrophiliclayers are envisaged in the context of the present application asdescribed below.

WO 2012/010767 describes a bag (or pouch or pocket) for forming animplantable artificial organ, which comprises a closed shell made of asemi-permeable membrane. This bag also comprises a sheet contained inthe shell, the sheet comprising projections (protuberances) on thesurface thereof for maintaining a space for cells between the sheet andthe shell.

US 20060067917 doesn't describe the membranes and encapsulating chamberdisclosed herein. The device of D2 is different from the devicesdisclosed herein, in its design, and can't be confused with theencapsulation chamber of the present application, as the membranes ofthe device of D2 are monolayer membranes (104, 106 and 112, of FIG. 1).

WO 2000/060051 describes an encapsulation chamber, the semi-permeablemembranes of which can be made from different materials and polymers(see page 21, line 15 to page 22, line 23 of this document). One shouldalso note that WO 2000/060051 envisages the use of various materialswithin the macroencapsulation device, in order to maintain the cells(page 21, line 30 to page 22, line 11).

However, there is a need to make available to surgeons novelbioartificial organs which exhibit, in particular, advantageousbiomechanical characteristics, i.e. good resistance after implantation.This is because bioartificial organs are intended to be implanted,generally in the intraperitoneal cavity or in the extraperitoneal spaceand are liable to undergo tensile or shear forces according to themovements of the recipient patient.

Moreover, these bioartificial organs must be able to contain a largenumber of cells, in order to be able to have a prolonged physiologicaleffect after implantation in the patient. It is therefore necessary todesign organs which are sufficiently large to do this, but they thenhave the drawback that they risk tearing after implantation due to thepatient's movements (this problem being less significant for microorganscontaining only a limited number of cells). Increasing the thickness ofthe membranes in order to improve the mechanical strength cannot be asolution since the diffusion of the molecules of interest is greatlyreduced when the thickness of the membrane increases.

It is therefore advisable to develop novel semi-permeable membranes withimproved mechanical properties for the manufacture of bioartificialorgans. The selective permeability properties must be at least retained.

In a first embodiment, the invention thus relates to a chamber forencapsulating secreting cells producing at least one substance oftherapeutic interest, comprising a closed shell made of a semi-permeablemembrane, delimiting a space capable of containing the secreting cellsproducing at least one substance of therapeutic interest, characterizedin that said membrane comprises at least one layer of porousbiocompatible polymer, and one layer of non-woven biocompatible polymer.

The documents cited above don't describe nor suggest such anencapsulation chamber for secreting cells, which comprises asemi-permeable membrane, which membrane comprises at least one layer ofporous biocompatible polymer, and another layer of non-wovenbiocompatible polymer.

Furthermore, as will be seen in examples (in particular examples 6 and7), the disclosed chambers show a higher mechanical resistance when usedin a bioartificial organ, in particular after in vivo implantation.

It is recalled that the term “biocompatible” is said of a material whichis well tolerated by a living organism and which does not cause arejection reaction, a toxic reaction, a lesion or a harmful effect onthe biological functions of the latter. This does not exclude thepossibility of an inflammatory reaction due to the insertion of thematerial into the organism or of an immune reaction in the case of abiocompatible organ comprising exogenous cells; this immune reaction isnot therefore due to the organ as such, but instead due to its content(secretion of chemokines by the exogenous cells).

As seen above, the semi-permeable membrane has a cut-off threshold, themolecules having a weight above this cut-off threshold being unable tocross the membrane, while the molecules having a weight below thiscut-off threshold can cross the membrane. The determination of thecut-off threshold is carried out by those skilled in the art accordingto the characteristics of the molecules that they wish to stop or allowto penetrate.

In one preferred embodiment, and in order to allow the passing of smallmolecules such as insulin, glucagon or glucose and to stop the effectormolecules of the immune system (such as cytokines), this cut-offthreshold is between 100 kDa and 500 kDa, more preferably between 100kDa and 150 kDa.

The internal diameter of the pores of the porous polymer makes itpossible to obtain the desired cut-off threshold. Thus, in oneparticular case, the internal diameter of the pores present on the layerof porous biocompatible polymer is between 5 and 100 nm, entirelypreferably between 5 and 50 nm.

Non-woven polymer

It is recalled that a non-woven polymer (non-woven) is such that thefibres thereof are maintained randomly. It is thus a sheet consisting offibres oriented in a particular direction or randomly, bonded byfriction and/or cohesion and/or adhesion. The fibres are thus arrangedstatistically, i.e. deposited randomly. Consequently and due to therandom arrangement of the fibers, the non-woven polymer is permeable tosubstances, and there is no control of the size of the substances thatcan diffuse within the non-woven polymer.

Non-woven polymers can be produced using polymeric fibres of any type.Mention may thus be made of polyesters: PET (poly(ethyleneterephthalate)), PBT (poly(butylene terephthalate)), PVC (poly(vinylchloride)), PP (polypropylene), PE (polyethylene) or blends of thesepolymers.

Polyamides or polycarbonates can also be used to produce non-wovenpolymers.

Preferably, the non-woven polymer is chosen from polycarbonate (PC),polyester, polyethyleneimine, polypropylene (PP), poly(ethyleneterephthalate) (PET), poly(vinyl chloride) (PVC), polyamide andpolyethylene (PE). Blends of these polymers can also be used forproducing the non-woven polymer. Poly(ethylene terephthalate) (PET) isparticularly preferred.

Generally, this non-woven polymer is obtained by the meltblown method.The composition thereof is an entanglement of microfibres which havebeen “melt blown”.

This method of production is particularly suitable for polymers whichcan be melt spun, in particular polypropylene, poly(ethyleneterephthalate), polyamides or polyethylene.

This method generates non-wovens of greater mechanical strength.

In one particular embodiment, said membrane comprises two layers ofporous biocompatible polymers, on either side of the layer ofbiocompatible non-woven polymer. Thus, this layer of biocompatiblenon-woven polymer is located, positioned or situated between these twolayers of porous biocompatible polymers.

Such an embodiment makes it possible to optimize the strength of thedevice. Indeed, this layer of non-woven can be considered to behave likea “sponge”, which gives it the capacity to absorb impacts and to deform,thus increasing the rigidity of the membrane in situ, but which canprove to be troublesome in the presence of cells, which can have atendency to form aggregates around this non-woven. Locating the layer ofnon-woven between two porous layers of biocompatible polymers thus makesit possible to prevent the aggregation of cells while at the same timeproviding the device with additional protection/strength, and with noeffect on the molecular diffusion of the biological substances.

It is not necessary for the porous and non-woven biopolymers to beidentical.

Likewise, in the presence of two layers of porous biopolymers, thelatter can be the same polymer or different polymers.

Porous Biocompatible Polymer

The porous biocompatible polymer consists of a polymer known in the art.Thus, it may be chosen from polycarbonate (PC), polyester,polyethyleneimine, polypropylene (PP), poly(ethylene terephthalate)(PET), poly(vinyl chloride) (PVC), polyamide and polyethylene (PE).

In one particular embodiment, at least one layer or the two layers, asappropriate, is (are) made of poly(ethylene terephthalate) (PET).

The pore formation is carried out by any method known in the art. Inparticular, it is possible to use the electron bombardment method or theheavy ion bombardment method (this second technique is in particulardescribed in patent U.S. Pat. No. 4,956,219). In the case of heavy ionbombardment, the density of the heavy ions bombarded at the surface ofthe biocompatible support determines the pore density, while thechemical erosion treatment time determines the pore size.

The membranes are thus prepared using the “track-etching” process knownin the prior art and described in particular in patents U.S. Pat. No.4,956,219, DE19536033 or CH701975.

This technology consists in irradiating a polymer film by means ofenergetic heavy ions, resulting in the formation of linear latent tracescharacterized by a local degradation of this polymer; these traces arethen revealed in the form of pores by means of a selective chemicalattack.

The membrane is irradiated with a beam of heavy ions. The heavy ionspass through the entire thickness of the polymer film. In passingthrough the polymer, the heavy ions destroy or cut the polymer chainsand thus form a clean straight opening through the material. The finalalignment of the pores is determined by the angle of the beam relativeto the polymer film during the irradiation process. The beam may thus beperpendicular to the polymer film or at any other angle.

In the next step, the film is passed through a bath of a strong acidsuch as nitric acid and the openings become pores after contact withalkaline solutions such as sodium hydroxide or potassium hydroxide.

Contrary to the rest of the film, these openings made by the ions allowthe alkaline solution to pass through, said alkaline solution fillingthem and allowing the etching of the pores by removing the material(polymer) around these openings.

The pore size is controlled by the concentration of the alkalinesolution, the contact time and the temperature of the solution.

If polyester or polycarbonate is used, the membrane obtained ishydrophilic and can either be used as it is or else be treated usingsurface treatment processes (plasma, spraying or coating).

The preparation of membranes according to this “track etching”technology is more precisely described in patents U.S. Pat. No.4,956,219 and CH701975.

This technology enables the production of porous polymer membranescharacterized in particular by a flat surface and a narrow cut-offthreshold.

The advantage of using membranes obtained by this technology is thegreat accuracy of the pore size, of the number of pores, and of theshape of the pores.

The pores are preferentially cylindrical, but this technology can alsomake it possible to obtain pores of other shape, such as of conicalshape.

Preferentially, the pores are aligned, and have an angle of between 10°and 45°, relative to the vertical, but can also have angles >45° or<10°. These angles are obtained according to the angler of the beam ofions during the bombardment of the membrane.

This technology is applicable to various materials, such aspolycarbonate (PC), polyester (PET) or polyimide (PI). Polyamide,poly(vinylidene fluoride), polyacrylate or polyolefins can also be used.

This method makes it possible to easily obtain pores with a controlledsize of between 0.02 μm and 15 μm, a pore density of between 10³pores/cm² and 10¹⁰ pores/cm² and membranes with a thickness of between 5μm and 80 μm.

It is to be noted that, without the treatment to form pores on thebiocompatible polymer, such polymer would remain impervious to anysubstance, and would not allow diffusion of the substance of interestfrom the inner part of the biocompatible organ to the outer part. Thepores only allow the diffusion of substances that are below the cutoff(i.e. that are smaller than the pore diameter).

It is thus clear that the layer of the non-woven biocompatible polymerand layer of the porous biocompatible polymer are different layers, madeof different materials, and presenting different properties (inparticular with regards to the passing and diffusion of substancesthrough each layer).

In one preferred embodiment, at least one of the layers of porousbiocompatible polymer of the membrane is made hydrophilic. Thehydrophilicity property can be achieved by generating polar sites at thesurface of this layer of porous biocompatible polymer. This surfacemodification can be carried out by physical means (such as thegenerating of charged polar sites at the surface, in particular byplasma surface treatment, by corona discharge or by electromagneticdischarge at atmospheric pressure or under vacuum) or chemical means (analkaline treatment, in particular with sodium hydroxide, can beenvisaged).

Preferentially, the layer of porous biocompatible polymer is treatedwith a radiofrequency argon, hydrogen, oxygen or air plasma. It can betreated at a plasma reactor emission power of between 3 and 10 watts perliter of reactor capacity, for between approximately 1 and 20 minutes.The treatment can also be carried out using a microwave plasma, at thesame power, but for 5 seconds to 20 minutes. Preferably, the plasmatreatment is carried out under vacuum.

Patent applications WO 02/060409 and WO 2012/017337 describe inparticular the plasma surface treatment for introducing polar sites ontothe porous biocompatible polymer.

After at least one layer of porous biocompatible biopolymer has beenmade hydrophilic, it is possible to cover it with at least one layer ofhydrophilic polymer, or even with two layers of different hydrophilicpolymers. An active molecule can optionally be contained in at least onelayer of hydrophilic polymer.

WO 02/060409 and WO 2012/017337 also describe the addition of at leastone hydrophilic polymer on the surface of a porous biocompatiblepolymer, said surface having been treated to make it hydrophilic, inparticular by adding polar sites.

Hydrophilic Polymer

For the purposes of the invention, what constitutes a hydrophilicpolymer is a polymer or a blend of polymers, which, after application ona film of porous biocompatible polymer, has an angle value of less than40°, preferably less than 30°, after measurement according to the“sessile drop” test described in Example 2 of WO 02/060409.

It should be noted that the angle value according to the “sessile drop”test can vary depending on the treatment of the polymer. Thus, contactangles of less than 20° (of about 16-17°) can be observed for thebiocompatible biopolymer, when two plasma treatments are carried out,this angle increasing (generally less than 30°) when the hydrophilicpolymer (in particular HPMC) is deposited after the two plasmatreatments. If a blend of hydrophilic polymers, which also contains amolecule with biological activity (in particular an HPMC,ethylcellulose+heparin mixture), is used, the angle may be greater than30°, but remains less than 40°.

Preferentially, the hydrophilic polymer is soluble in water. This isbecause, due to of the implantation of the bioartificial organ in thebody of a host organism, use of organic solvents is excluded since theircomplete elimination is difficult, and their presence, even in smallamounts, is not compatible with a therapeutic or surgical use in humansor animals.

Preferably, the hydrophilic polymer material is chosen from thefollowing hydrophilic polymers:

-   -   celluloses and derivatives thereof, such as ethylcellulose (EC),        hydroxypropylmethylcellulose (HPMC) or carboxymethylcellulose        (CMC);    -   polyacrylamides and copolymers thereof;    -   polyvinylpyrrolidone (PVP) and copolymers thereof;    -   polyvinyl alcohols;    -   vinyl acetate copolymers, such as a poly(vinyl        acetate)/poly(vinyl alcohol) copolymer;    -   polyethylene glycols;    -   propylene glycols;    -   hydrophilic poly(meth)acrylates;    -   polysaccharides;    -   chitosans.

As hydrophilic polymer, use is made of both a polymer materialconsisting of one of the hydrophilic polymers as defined above and ablend of several of the hydrophilic polymers above, generally a blend oftwo or three of the hydrophilic polymers above.

Preferably, the hydrophilic polymer is chosen from cellulose-basedcompounds, in particular HPMC, EC, TEC or CMC, polyvinylpyrrolidones,poly(vinyl alcohol)s, or polyacrylates such as poly(hydroxyethylacrylate) (HEA) or acrylic acid copolymers.

The hydrophilic polymer may also be composed of a blend of two or morehydrophilic polymers mentioned above, in particular a blend of HPMC andCMC, or of HPMC and EC.

Celluloses and cellulose derivatives, in particularhydroxypropylmethylcellulose (HPMC), are preferred.

Membrane Lamination

For greater mechanical stability, the porous biocompatible polymermembrane is reinforced using a membrane made of non-woven.

The combining of a non-woven polymer and of the porous membrane ofbiocompatible polymer is preferentially carried out by lamination, usingmethods known in the art, such as thermal lamination, with or withoutthe presence of adhesives, preferably without adhesive.

Thus, the reinforcement of the membrane can be improved via a multilayersystem alternating layers of woven or non-woven polymers and ofbiocompatible porous polymers. However, any degradation of the diffusionproperties should be avoided.

In particular, the mechanical stability can be increased by combining athin functional membrane which has a high pore density with a thickprotective membrane which has a low pore density.

There is no limitation to the number of layers of polymers that can beused to manufacture the membrane.

Active Molecule

As indicated above, the hydrophilic polymer deposited on the layer ofporous biocompatible polymer can optionally contain an active molecule.

This “active molecule” is mixed with the hydrophilic polymer. It isintended to be released into the medium surrounding the semi-permeablemembrane in particular in order to reduce the inflammation due to theimplantation of the bioartificial organ, and/or to induce a positiveresponse (in particular increased vascularization) by the tissue(s) ofthe patient receiving the bioartificial organ.

Thus, the active molecule is chosen from anti-inflammatory agents,anti-infective agents, anaesthetics, growth factors, agents whichstimulate angiogenesis and/or which induce vascularization, agents whichinduce healing, immunosuppressive agents, antithrombotics includingantiaggregants and anticoagulants, angiotensin-converting enzyme (ACE)inhibitors, or any molecule which stimulates insulin secretion (IGF,glucagon-like peptide 1 (GLP-1) or its derivatives, incretin mimetics).

Among the anti-inflammatory agents, mention may be made of non-steroidalanti-inflammatories (NSAIDs), such as acetaminophen, aminosalicylicacid, aspirin, celecoxib, choline magnesium trisalicylate, declofenac,diflunisal, etodolac, flurbiprofen, ibuprofen, indomethacin, interleukinIL-10, IL-6 mutein, anti-IL-6, NO synthase inhibitors (for example,L-NAME or L-NMDA), interferon, ketoprofen, ketorolac, leflunomide,mefenamic acid, mycophenolic acid, mizoribine, nabumetone, naproxen,oxaprozin, pyroxicam, rofecoxib, salsalate, sulindac and tolmetin, andcorticoids such as cortisone, hydrocortisone, methylprednisolone,prednisone, prednisolone, betamethasone, betamethasone dipropionate,betamethasone valerate, beclomethasone dipropionate, budesonide,dexamethasone sodium phosphate, flunisolide, fluticasone propionate,paclitaxel, tacrolimus, tranilast, triamcinolone acetonide, fluocinoloneacetonide, fluocinonide, desonide, desoximetasone, fluocinolone,triamcinolone, clobetasol propionate and dexamethasone. Ibuprofen isparticularly suitable and preferred.

Use is preferably made of antithrombotics such as antiaggregants(acetylsalicylic acid, clopidogrel, ticlopidine, dipyridamole,abciximab, eptifibatide and tirofiban), anticoagulants (heparin,bivalirudin, dabigatran, lepirudin, fondaparinux, rivaroxaban,epoprostenol, warfarin, phenprocoumon, protein C, drotrecogin alfa,antithrombin, pentosan) and thrombolytics (alteplase, urokinase,tenecteplase and reteplase).

The use of a heparin is particularly preferred.

In another embodiment, ibuprofen is used.

In addition, it is possible to use a molecule which makes it possible toinduce vascularization of the tissues surrounding the bioartificialorgan, in particular PDGF (platelet derived growth factor), BMP (bonemorphogenetic protein), VEGF (vascular endothelial growth factor), VPF(vascular permeability factor), EGF (epidermal growth factor), TGF(transforming growth factor) and FGF (fibroblast growth factor).

It is also possible to use IGF-1 and IGF-2, a neurotrophic factor (NGF).

In one particular embodiment, a cell growth factor is chosen whichpromotes vascularization by inducing angiogenesis, such as basicfibroblast growth factor (bFGF), vascular endothelial growth factor(VEGF), platelet derived endothelial cell growth factor (PDGF A or B),bone morphogenetic protein (BMP 2 or 4), or hepatocyte growth factor(HGF).

For the preparation of the layer of hydrophilic polymer and biologicallyactive molecule, the hydrophilic polymer or the blend of hydrophilicpolymers is dissolved in water.

The addition of the hydrophilic polymer optionally containing an activemolecule to the layer of porous biocompatible polymer is carried outaccording to the methods described in WO 02/060409 and WO 2012/017337.

In another embodiment, it is possible to add, at the surface of theporous biocompatible polymer, two layers each comprising a hydrophilicpolymer and at least one biologically active molecule, as described inWO 2012/017337.

Physical Characteristics of the Biocompatible Membrane

In the preferred embodiment, the membrane according to the inventioncomprises two layers of porous biocompatible polymer, each covered withat least one hydrophilic polymer, which surround the layer of non-woven.

Pore Diameter and Density

As seen above, the pores are introduced into each of the layers ofporous biocompatible polymer using methods known in the art. It ispreferred for at least the layer (if it is the only one) or one of thetwo layers of porous biocompatible polymers to have a pore densitygreater than 10⁶ pores/cm², preferably greater than 10⁷ pores/cm². Thispore density is generally less than 10¹¹ pores/cm², preferably less than10¹⁰ pores/cm². Use is therefore made of membranes which can have a poredensity preferentially greater than 10⁶ pores/cm², more preferablygreater than 10⁷ pores/cm². This density is preferentially less than10¹¹ pores/cm², or even less than 10¹⁰ pores/cm². This density istherefore between 10⁶ pores/cm² and 10¹¹ pores/cm². A density greaterthan 10⁹ and less than 10¹⁰ pores/cm² is perfectly suitable.

As seen above, the pores of the layers of porous biocompatible polymerhave an internal diameter such that they allow semi-permeability of themembrane.

Thus, at least one of the two layers (or the only layer if such is thecase) of porous biocompatible polymer has pores which have an internaldiameter greater than 5 and preferably greater than 10 nm, and less than100 nm, and preferably greater than 10 nm and less than 50 nm, morepreferably less than 40 nm. A pore diameter of less than 90 nm is alsovery favorable for this layer of porous biocompatible polymer, as suchpore diameter maintains the semi-permeability property, that is soughtfor the membrane. The pore density is then advantageously greater than2.10⁹ and less than 4.10¹⁰ pores/cm².

When the membrane has two layers of porous biocompatible polymers, theinternal diameter of the pores of one of the layers is preferentially asabove.

The internal diameter of the pores of the second layer may be larger,the cut-off effect at the desired size being given by the diameter ofthe pores of the first layer. Thus, the internal diameter of the poresof the second layer may be greater than 100 and less than 2000 nm,preferably greater than 200 nm. These pores preferably have an internaldiameter less than 1000 nm. An internal pore diameter greater than 400and less than 600 nm, or of approximately 500 nm, is perfectly suitable.The pore density is then advantageously greater than 5.10⁶ and less than5.10⁷ pores/cm².

When the membrane comprises two layers of porous biocompatible polymer,which surround the layer of non-woven, it is preferable for theencapsulating chamber to be such that the layer for which the porediameter is the smallest is situated inside the chamber (in contact withthe secreting cells producing at least one substance of therapeuticinterest) and that the layer for which the pore diameter is the widestis situated on the outside (in contact with the patient's body).

Membrane Thickness

In one preferred embodiment, the total thickness of the membrane(comprising the layer of non-woven polymer and the layer(s) of porouspolymer(s)) is greater than 45 μm. It is generally, and preferably, lessthan 200 μm, but can also be greater than this size; thicknesses rangingup to 300 μm, or even beyond, can in particular be envisaged.Preferably, it is greater than 50 μm. It is also preferentially lessthan 150 μm. This membrane thus generally has a thickness of between 45and 200 μm.

When the membrane has two layers of porous biocompatible polymers, saidlayers can have the same thickness or have different thicknesses.

The layer of non-woven polymer generally has a thickness greater than 40μm, preferably greater than 60 μm, more preferably greater than 80 μm.This layer has a thickness generally less than 250 μm and preferablyless than 150 μm. Thus, the thickness of the layer of non-woven polymeris often between 40 μm and 150 μm.

When the membrane has only one layer of biocompatible polymer, saidlayer then has a thickness greater than 5 μm. This layer is less than200 μm, preferably less than 100 μm, being, however, preferably lessthan 50 μm.

When the membrane has two layers of porous biocompatible polymer, andsaid layers have different thicknesses, the thickness of the first layeris then greater than 5 μm. It is also preferably less than 200 μm, butpreferably less than 40 μm; a thickness less than 15 μm (and preferablygreater than 5 μm) is perfectly suitable. This thickness ispreferentially the thickness of the layer which has pores with thesmallest internal diameter, if the internal pore diameter is differentfor the two layers.

The thickness of the second layer is generally greater than 25 μm. It ispreferably less than 200 μm, preferably less than 100 μm, morepreferably less than 50 μm; a thickness of between 30 and 50 μm isperfectly suitable.

The thickness of each layer of hydrophilic polymer optionally present onone or the two layer(s) of porous biocompatible polymers is negligible,compared with the total thickness of the membrane. It is in factpreferably less than 500 nm and generally between 25 and 250 nm.

In one preferred embodiment, the membrane has two layers of porousbiocompatible polymers on either side of a layer of non-woven polymer.

In this embodiment, one layer of porous biocompatible polymer has poreswith an internal diameter greater than 100 nm, preferably greater than200 nm, more preferably greater than 400 and less than 1000 nm, morepreferably less than 600 nm, preferably at a density of about 5.10⁷pores/cm². It is then advantageous for this layer to be the one with athickness of between 25 and 200 μm (see above).

The other layer of porous biocompatible polymer has pores with aninternal diameter greater than 5 nm, preferably greater than 10 nm (andgenerally less than 100 nm, preferably less than 50 nm, preferably lessthan 40 nm), preferably at a density of about greater than 2.10⁹pores/cm². This density is also preferentially less than 7.10⁹pores/cm².

It is advantageous for this to be the layer with a thickness of between5 and 200 μm (preferably 5 to 15 μm).

Encapsulation Chamber

The invention also relates to a chamber for encapsulating secretingcells producing at least one substance of therapeutic interest,comprising a closed shell made of a membrane according to the invention,delimiting a space capable of containing the secreting cells producingat least one substance of therapeutic interest. This encapsulatingchamber can also be referred to as a “pouch” and makes it possible toform a bioartificial organ which is implantable in the patient.

In one particular embodiment, this encapsulating chamber also comprisesa biocompatible sheet contained in said shell, said sheet preferablycomprising projections (also designated as protuberances) at itssurface. These projections are advantageous for maintaining a space forthe cells between the sheet and the shell, but also for distributing thecells in a homogeneous and planar manner, thus making it possible tomaximize the exchange surface. This sheet is preferentially made ofsilicone.

Such an embodiment is described in application WO 2012/010767. Thus, inone preferred embodiment, the shell is formed from two membranes whichare heat-welded together. Use may be made of the method described in WO2012/010767 or a method of heat-welding using ultrasound, known in theart. The method for forming the shell is simple and makes it possible toenclose the sheet in the shell.

Shape of the Chamber

In one preferred embodiment, the encapsulating chamber is circular. Sucha shape has several advantages:

-   -   absence of “corners” or protruding parts which are capable of        creating cell or inflammatory aggregates during the        implantation,    -   ease of manufacture of the encapsulating chamber (no need to        orient the two membranes and the sheet before the heat-welding).

In one particular embodiment, the diameter of the encapsulating chamberis greater than 3 cm, preferably greater than 5 cm, or than 8 cm. It isgenerally less than 20 cm, and is preferentially less than 15 cm, orthan 14 cm. A diameter of between 8 and 14 cm is perfectly acceptable.

When the chamber is not round, the largest dimension thereof isgenerally greater than 3 cm, preferably greater than 5 cm, or than 8 cm.It is generally less than 20 cm, and is preferentially less than 15 cm,or than 14 cm.

Volume of the Chamber

As seen above, the encapsulating chamber preferentially makes itpossible to manufacture a “macro” organ when the secreting cellsproducing at least one substance of therapeutic interest are introducedtherein, i.e. it allows said cells to secrete this substance for aconsiderable period of time (greater than 3 months, preferentiallygreater than 6 months) at levels which are of physiological interest(i.e. making it possible to meet the patient's need). The encapsulatingchamber should therefore be able to receive a large number of cells.

It is generally estimated that the preferred internal volume of theencapsulating chamber should be greater than 15 ml, preferably greaterthan 20 ml, more preferably greater than 25 ml, and can rise to 50 ml,for use in humans. For use in other animals, the volume will bedifferent (approximately 1 ml in rats, for example).

Such encapsulating chambers must be able to contain a large number ofcells. In the context of the treatment of diabetes, it must be possibleto encapsulate the equivalent of at least 500 000 islets of Langerhans,preferably the equivalent of more than 700 000 islets, and optionally upto the equivalent of one million islets of Langerhans. In the knowledgethat one islet contains, on average, about 1000 cells, this gives anestimation of the number of cells that the encapsulating chamberaccording to the invention can contain.

The number of cells will obviously vary according to the type of cellsthat it is desired to encapsulate and implant in the patient.

In one preferred embodiment, the membrane forming the encapsulatingchamber comprises two layers of porous biocompatible polymers on eitherside of the non-woven polymer. In this embodiment, it is preferred forat least the internal layer (situated inside the chamber after formationof the chamber) to be the layer on which the pores provide thesemi-permeable nature of the membrane (cut-off threshold), i.e. whichhas the pores that have an internal diameter greater than 5 nm (andgenerally less than 100 nm) or having the other dimensions mentionedabove.

The layer external to the shell (in contact with the patient's tissuesand cells) can have pores with a larger internal diameter, in particulargreater than 100 nm, but preferably less than 2000 nm, or having theother dimensions mentioned above.

In one embodiment, and as described in WO 2012/010767, the encapsulatingchamber can comprise at least one connector (in particular attached tothe shell and/or the sheet), which makes it possible to establish acommunication between the exterior and the interior of the shell.Connecting these connectors to flexible tubes makes it possible to filland empty the chamber.

Bioartificial Organ

The invention thus relates to a bioartificial organ comprising at leastone encapsulating chamber according to the invention. Such abioartificial organ also advantageously presents the tubes connected tothe connectors and making it possible to fill and empty thebioartificial organ, making it possible to renew the content of thebioartificial organ when it is implanted in a patient, withoutperforming an explantation.

This bioartificial organ may contain various cell types.

Cells Encapsulated in the Bioartificial Organ

The cells present in the bioartificial organ produce at least onebiologically active substance of interest. They can in particular beinsulin-secreting cells or islets of Langerhans, which produce insulin,when the encapsulating chamber is intended for the manufacture of abioartificial pancreas.

The cells may also be hepatic cells when the encapsulating chamber isintended for the manufacture of a bioartificial liver.

In one particular embodiment, the cells are transfected or transformedwith at least one nucleic acid allowing the expression of a biologicallyactive substance of interest. Among the biologically active substancesof interest, mention may be made, by way of illustration, of insulin,cytokines, peptide hormones, growth hormone, coagulation factors VIIIand IX and calcitonin.

Generally, the term “biologically active substance” is intended to meana substance which is released or secreted by the cell which produces itand which exerts its effect on a target cell or a target molecule in thehost organism, for instance a neurotransmitter, a hormone, a growthfactor, a coagulation factor or a cytokine.

A great diversity of cells can be used, including immortalized celllines, for instance primary cultures of dividing cells, or elsepluripotent stem cells.

The cells can, for example, be myoblasts, which are cells that areprecursors of muscle cells derived from the stem cell populations of themesoderm, and which can be easily transformed with a nucleic acidallowing the expression of the biologically active substance ofinterest. Those skilled in the art may advantageously refer, forexample, to WO 94/02129, WO 93/03768 or WO 90/15863.

Preferably, the cells contained in an encapsulating chamber according tothe invention are embedded in a matrix, such as a matrix of collagentype IV or of fibrin, where appropriate in combination with laminin,entactin and heparan sulphate.

The cells contained in an encapsulating chamber according to theinvention can generally be embedded in a matrix composed of any productor combination of products allowing the immobilization of these cells ina viable form.

The cells producing at least one biologically active substance ofinterest can also be encapsulated in an alginate matrix.

Manufacture of an Encapsulating Chamber

The encapsulating chamber is manufactured by any method known in theart.

Use is preferably made of the teaching of WO 2012/010767, which shouldbe considered to be an integral part of the present application.

The invention thus relates to a method for manufacturing anencapsulating chamber according to the invention, comprising a step ofheat-welding two membranes according to the invention (or even a foldedmembrane), in such a way as to form a pouch intended to receive cellsproducing at least one biologically active substance of interest.

In one particular embodiment, as seen above, the encapsulating chambercontains a sheet, and also one or more connectors. The method formanufacturing such a pouch is described in WO 2012/010767. The reader isinvited to refer to WO 2012/010767 for more detailed explanationsregarding the process for manufacturing the encapsulating chamber.

DESCRIPTION OF THE FIGURES

FIG. 1: Permeability of poly(ethylene terephthalate) (PET) orpolycarbonate (PC) membranes according to the invention, treated or nottreated with heparin, ethylcellulose (EC) andhydroxypropylmethylcellulose (HPMC), to glucose (A), insulin (B) andIgGs (C) under static conditions.

FIG. 2: Insulin secretion by rat pancreatic islets stimulated withglucose through a PET membrane according to the invention, treated ornot treated with heparin, EC and HPMC. A beginning of diffusion of theinsulin starting from 4 hours and a permeability which appears to beimproved at 24 hours by the surface treatment are observed.

FIG. 3: Images of the sections prepared 30 days after the implantationof poly(ethylene terephthalate) (PET) or polycarbonate (PC) membranesaccording to the invention, treated or not treated with heparin, EC andHPMC. The surface treatment decreases fibrosis and cell infiltration(black arrows) and increases vascularization (*) for the two types ofmembrane.

FIG. 4: Appearance of the bioartificial organs after 15 days ofimplantation in pigs. One of the devices is composed of monolayer PCmembranes and the other of multilayer PET membranes. The device with PCmembranes shows wide tears. The device with multilayer PET membranesdoes not, for its part, show any macroscopic damage. Said multilayer PETmembranes were thus analyzed by scanning electron microscopy, whichdemonstrated no microcracks.

EXAMPLES Example 1 Manufacture of Semi-Permeable Membranes

The membranes are manufactured such that two porous PET (poly(ethyleneterephthalate)) layers were prepared from biocompatible PET films by the“track-etching” process, followed by lamination with the layer ofnon-woven PET having a density between 30 and 60 g/m2 (situated betweenthe two porous biocompatible PET layers). A thermal lamination iscarried out without the use of adhesives. One of the porous PET layershas a pore density between 2.10⁹ and 7.10⁹ pores/cm² with an internalpore diameter between 10 and 30 nm. The thickness of this membrane isbetween 8 and 12 μm. The other porous PET layer has a pore densitybetween 10⁷ and 5.10⁷ pores/cm² with an internal pore diameter between400 and 600 nm. The thickness of this membrane is between 30 and 50 μm.The total thickness of the membrane is less than 200 μm.

Example 2 Surface Treatment of the Membranes

The membranes prepared according to Example 1 were subjected to asurface treatment according to the protocol of Example 1 of WO2012/017337.

The membranes are functionalized with a first layer of heparin mixedwith a solution of ethylcellulose (EC), then covered with a layer ofhydroxypropylmethyl-cellulose (HPMC).

Example 3 Characterization of the Membrane Permeability

Tests for glucose-permeability, insulin-permeability and immunoglobulin(IgG)-permeability of the previously prepared membranes were carried outaccording to the following protocol:

Material

Diffusion chamber consisting of a top compartment and a bottomcompartment separated by the membrane, the permeability of which it isdesired to test (the leaktightness between the two compartments isprovided by a seal), glucose (Fischer Scientific, Illkirch, France, ref:G/0500/53), NaCl, IgG (Sigma, Lyon, France, ref: 19640), insulin (Sigma,ref: 19278), distilled water.

Preparation of Solutions

-   -   Physiological Saline

-   For 1 l: 9 g of NaCl are dissolved in 1 l of distilled water.    -   Glucose (4 g/l)

-   For 1 l: 4 g of glucose are dissolved in 1 l of physiological    saline.    -   IgG (5.75 μg/ml)

-   For 60 ml: 34.5 μl of stock solution of IgG (10 mg/ml) are diluted    in 59.966 ml of physiological saline.    -   Insulin (100 μg/ml)

-   For 60 ml: 60 μl of stock solution of insulin (10 mg/ml) are diluted    in 59.960 ml of physiological saline.    Protocol

3 ml of physiological saline are introduced into the bottom compartmentof the diffusion chamber, and the membrane, the permeability of which itis desired to test, is placed on the physiological saline while avoidingthe presence of air bubbles. 3 ml of glucose solution are introducedinto the top compartment, then the diffusion chamber is closed withparafilm and is incubated at 37° C.

At the end of the incubation time, 1 ml of the solution contained in thetop compartment of the diffusion chamber is removed after gentlehomogenization. The membrane is then removed and 1 ml of the solution ofthe bottom compartment is removed after homogenization.

Enzymatic assaying of the glucose is carried out using the Glucose RTU®kit (BioMérieux, Craponne, France ref: 61 269). The insulin and the IgGsare assayed using the bicinchonic acid (BCA) method by means of theQuantipro BCA Assay kit (Sigma, ref: QPBCA-1KT). The results areexpressed as percentage permeability, calculated in the following way:Permeability (as%)=(C_(bottom compartment)/C_(top compartment)+C_(bottom compartment))×100

C: concentration of glucose, IgG or insulin.

At equilibrium, the concentrations in the top compartment and in thebottom compartment are identical, which corresponds to a maximumpermeability of 50%.

Results

The results are shown in FIG. 1. Multilayer poly(ethylene terephthalate)(PET) membranes according to the invention (Example 1), and also priorart membranes as described in WO 02/060409 or WO 2012/017337, made ofpolycarbonate and having a layer of heparin mixed with EC and a layer ofHPMC, were tested.

A slower diffusion of insulin and of glucose was observed with the PETmembranes. Without wishing to be bound by this theory, it is possiblethat this is due to the presence of the multilayers of which they arecomposed.

The PET membranes are totally impermeable to IgGs.

Example 4 Semi-Permeable Membrane Implantation Tests

The membranes are implanted in the peritoneal cavity of healthy Wistarrats, according to the protocol described in Example 3 of WO2012/017337.

The protocol relating to the taking of the samples was however modifiedand the samples are taken in the following way:

Taking Tissue Samples

Solutions Used

-   -   2.5% glutaraldehyde prepared, under a hood, from 25%        glutaraldehyde (Sigma, ref: G5882-10×10 ml) diluted to ten-fold        in ultrapure water.    -   PBS (reference: Gibco-14190-094).    -   Pot prefilled with 4% paraformaldehyde (Labonord, ref:        PFFOR0060AF59001).

The membranes tested are the PET membranes according to the invention(multilayer) and the PC membranes of the prior art, optionally havingundergone a surface treatment in order to deposit heparin, EC and HPMC.

The results are shown in FIG. 3: it is observed that the surfacetreatment with heparin reduces fibrosis and cell infiltration (blackarrows) and increases vascularization (*) for the two types of membrane.

Example 5 Test for Glucose-Stimulation of Islets through the Membrane

a) Isolation of Rat Pancreatic Islets

Animals Used

The animals used are male Wistar rats weighing 250-300 g (JanvierLaboratory, Le Gènes St. {circumflex over (l)}le, France). The rats arehoused in standard collective cages at a temperature of 23±1° C., and ahygrometry of 55±3% and with a cycle of 12 h of light and 12 h in thedark. SAFE-A04 feed (Villemoisson-sur-Orge, France) and water areavailable ad libitum. The animal experiments are carried out inaccordance with European directive 2010/63/EU.

Removal of the Pancreas

The animal is anaesthetized with a mixture of Imalgene 1000® (activeingredient: ketamine, Centravet ref: IMA004) supplemented with 2.7 ml ofRompun® (active ingredient: xylazin at 2%, Centravet ref: ROM001)injected intraperitoneally at a dose of 100 μl/100 g of body weight.

After having verified the absence of reflexes of the animal, the latteris laid on its back. A laparotomy is then performed and the bile duct isligatured at its duodenal opening. It is then catheterized at itshepatic opening and the animal is sacrificed by exsanguination. 10 ml ofcollagenase type XI (Sigma, ref: C7657) at 1 mg/ml at 4° C. are theninjected into the pancreas by means of the catheter.

The pancreas is then removed and placed in a 50 ml Falcon tubecontaining 3.75 ml of sterile “perfusion solution”. This solution iscomposed of 500 ml of HBSS (Hanks Balanced Salt Solution, Lonza, ref:BE10-527F), 2.1 ml of 8.4% sodium bicarbonate, 1.175 ml of 1M calciumchloride and 12.5 ml of 1M HEPES. In order to limit the action of theenzyme during the removal, the tubes containing the pancreases are keptin ice.

Digestion

Immediately after the pancreases have been removed, the tube is placedin a waterbath at 37° C. for 10 minutes. It is then vigorously stirredfor a few seconds in order for the tissue to be well dissociated. It isthen made up with a cold washing solution. The washing solution iscomposed of M199 (Sigma, ref: M0393-50L) supplemented with 0.35 g/l ofsodium bicarbonate (Sigma, ref: S-5761), with 10% of foetal calf serum(FCS, Lonza, ref: DE14-801F) and with 1% of anti-mycotic antibiotic(AMAB, Fisher, ref: W3473M).

The content of the tube is filtered on inserts (Corning Netwell inserts,Sigma, ref: CLS3480) and the filtrate is transferred into a 200 mlCorning tube which is centrifuged for 1 minute at 1200 rpm at 4° C. Thesupernatant is then removed and the pellet is resuspended with coldwashing solution, then transferred into a 50 ml Falcon tube. Aftercentrifugation for 1 minute at 1200 rpm at 4° C., a maximum amount ofsupernatant is removed before going on to the purification step.

Purification

The purification of the islets is carried out using a discontinuousgradient of Ficoll (Fisher, ref: BP525-500) which is composed of 3solutions of different densities prepared in the laboratory: 1.108(Ficoll 1): 1.108, 1.096 (Ficoll 2): 1.096 and 1.069 (Ficoll 3): 1.069.

The cell pellet is resuspended in 12 ml of Ficoll 1, and 10 ml of Ficoll2 then of Ficoll 3 are carefully added on the top. Finally, 5 ml of PBS(Fisher, ref: 20012-019) are deposited on the Ficoll 3. The wholeassembly is centrifuged for 4 minutes at 400 rpm at 4° C. and then for12 minutes at 2000 rpm at 4° C. The braking and accelerating speeds ofthe centrifuge are adjusted to the minimum so as not to disturb thegradients.

The islets are recovered at the interface between the Ficoll 2 and theFicoll 3, and are then washed three times in a cold washing solution inorder to remove any trace of Ficoll.

Culturing

The islets are cultured in M199 medium (Gibco, ref: 23340-020)containing 10% of FCS (Lonza, ref: DE14-801F) and 1% of AMAB (Fisher,ref: W3473M) in untreated 25 cm2 flasks (Dutscher, ref: 690195), for 24hours at 37° C. and in a humid atmosphere at 5% CO2.

b) Stimulation Test

Ten rat islets are placed in inserts (type of cylinders) at one end ofwhich the PET membrane is attached. This membrane is oriented in such away that the nanoporous membrane (which has pores with an internaldiameter between 10 and 50 nm and which is selective for molecules up to150 kDa) is on the inside of the insert, in contact with the rat islets,the layer which has the pores with a diameter of between 400 and 600 nmbeing oriented towards the outside of the insert.

The insert contains 400 μl of Krebs solution containing 10% of FCS and2.5 mM of glucose. The inserts thus filled are placed in wells of a24-well plate containing 1 ml of Krebs solution containing 10% of FCSand 25 mM of glucose. The 24-well plate is then incubated at 37° C. andsamples of medium contained in the wells are taken at 1 h, 2 h, 4 h, 6h, 8 h and 24 h. The insulin is then assayed in the samples using theELISA method (Mercodia, ref: 1250-01).

The islets are also sampled and placed in 50 μl of lysis buffer(ThermoScientific, ref: 78501), supplemented with a protease inhibitor(ThermoScientific, ref: 78441), in order to extract the total proteins.The extraction is carried out by placing the tubes on ice for 30 min,while regularly vortexing the samples. The total protein content of theislets is determined by means of a Bradford assay and is used tonormalize the secretion of insulin between the various islet cultures.

Example 6 Implantation and Explantation of MAILPAN® in Pigs

An encapsulating chamber (MAILPAN®, for MAcro-encapsulation d'ILotsPANcréatiques [macro-encapsulation of pancreatic islets]) is preparedaccording to the method described in WO 2012/010767. Two semi-permeablemembranes are welded together. This encapsulating chamber has aninternal sheet, and also connectors.

Anaesthesia

Premedication is systematic before any anesthesia and consists of theintramuscular administration of a combination of a butyrophenone: 2mg/kg azaperone (Stresnil*) and of 10 mg/kg ketamine (Imalgene*).

General anesthesia is carried out according to the protocol describedhereinafter:

-   -   the animals are taken, premedicated, to the operating block and        placed on the operating table lying on their side.    -   A peripheral vein is catheterized (G 22) on one ear and its        permeability is ensured by rinsing with a 0.9% NaCl solution.    -   The induction is carried out by intravenous injection of a        hypnotic (5 mg/kg thiopental or 4 mg/kg propofol) and of a        curarising agent (0.1 mg/kg pancuronium). It is immediately        followed by orotracheal intubation (Portex Blue Line,        low-pressure balloon, calibre 6 for a subject weighing 25 to 35        kg) and by pulmonary ventilation using a semi-closed circular        system connected to a respirator operating in controlled        pressure mode. The ventilation (FiO₂=0.5 FiN₂O=0.5) is adjusted        so as to maintain E_(T)CO₂ between 35 and 45 mmHg. The        respirator is a latest-generation human apparatus (GE Avance*,        Aisys* or Aespire*) fitted with current flow rate, pressure and        volume controls.    -   The anaesthesia is maintained on inhalation mode with isoflurane        (fraction inspired=2 vol %) with a fresh gas flow rate of 2        l/min of a 50%/50% O₂/N₂O mixture serving as vector gas.    -   If it proves to be necessary, the administration of subsequent        doses of pancuronium provides optimum muscle relaxation under        coverage of deep inhalation anaesthesia (MAC of isoflurane in        pure O₂=1.15 vol % and MAC of N₂O=110 vol %.        MAILPAN® Implantation

After anesthesia of the animals, the abdomen of the animals sent tosleep is made aseptic using 70% ethanol and then betadine (taking carenot to cause hypothermia) and is shaved using a scalpel blade. Alongitudinal incision of approximately 10 to 15 centimeters of the skinand muscle planes as far as the peritoneum is made in the middle of thecleared zone. After a median laparotomy, the prototype is implantedextraperitoneally, after being filled with physiological saline, and isattached to the wall with thread (Vicryl 2/0). The two catheters of theMAILPAN (one being used for the filling and the other for the emptyingof the islets in the MAILPAN, in a period subsequent to theimplantation) are connected to two injection chambers placedsubcutaneously, before ligature of the peritoneum by sinusoidalmovement, using 4-0 suture thread.

At the end of the surgery, the wounds are infiltrated with Naropein, andfentanyl will be administered IV before the animal is woken up. Fentanylgranules are administered per-operatively with the food intake, in aproportion of 2 mg/kg.

MAILPAN® Explantation

The MAILPAN devices are explanted 15 days and 60 dayspost-transplantation under general anesthesia in order to evaluate themechanical strength of the MAILPAN, the sterility thereof and thebiocompatibility thereof (vascularization at the surface, absence ofinflammation, absence of fibrosis and of inflammation on the surroundingtissues). Thus, samples of tissues surrounding the MAILPAN are taken ateach explantation of the device for subsequent histological tests. Thepigs are sacrificed after each explantation by intravenous injection ofKCl.

The tissue samples are taken under the same conditions as for the rat(see Example 5: same solutions for tissues and membranes, same analysescarried out).

Example 7 Analyses of the Membranes by Scanning Electron Microscopy

After sampling, the membranes are rinsed in ultrapure water and fixedfor 24 to 48 h at 4° C. in glutaraldehyde (Sigma, ref: G5882) diluted to2.5%. The fixed membranes are then rinsed for 10 minutes in ultrapurewater.

The samples are then dehydrated using successive baths of ethanol: twobaths of 10 minutes in 50% ethanol, one bath of 25 minutes in 70%ethanol, then one bath of 10 minutes in 95% ethanol and, finally, twobaths of 10 minutes in 100% ethanol. In order to completely remove thetraces of water which might still be present in the samples, anincubation for 2 minutes is carried out in hexamethyldisilazane (HMDS)(Sigma, ref: 440191).

After drying in the open air, the samples are then adhesively bondedflat on blocks (Delta Microscopies, ref: 75220), using carbon-conductingadhesive (Delta Microscopies, ref: 76510).

Once the adhesive has solidified, the samples are metallized bydepositing a thin layer of gold-palladium, and then of carbon.

The observation is carried out on a field-effect scanning electronmicroscope (SEM) (Hitachi S800) (in vitro imaging platform of theNeurochemistry Centre of Strasbourg) at a voltage of 5 KV, which makesit possible to obtain good resolution without damaging the samples.

Results

It is observed that the device produced with PC membranes shows widetears (FIG. 4).

On the other hand, the device produced with the multilayer PET membranesdoes not, for its part, show any macroscopic damage. Said membranes werethus analyzed by scanning electron microscopy, which demonstrated nomicrocracks (FIG. 4).

It therefore appears that the membranes according to the invention allowa diffusion similar to that observed for the prior art membranes andclearly have the property of semi-permeability (blocking IgGs, and otherproteins of the immune system). These membranes exhibit much betterresistance when they are used in a bioartificial organ implanted invivo.

Further data has been obtained for tensile strength in vitro, for PETmembranes. The strength is slightly higher for a tri-layer membrane (twoporous PET membranes surrounding a non-woven PET membrane), than for atwo-layer membrane (one porous PET membrane laminated on a non-woven PETmembrane). The tensile strength of the two-layer membrane is higher thanthe one for a mono-layer porous PET membrane.

The invention claimed is:
 1. A chamber for encapsulating secreting cells producing at least one substance of therapeutic interest, comprising a closed shell made of a semi-permeable membrane, delimiting a space capable of containing said secreting cells producing at least one substance of therapeutic interest, wherein said membrane comprises a layer of biocompatible non-woven polymer located between two layers of porous biocompatible polymers.
 2. The encapsulating chamber according to claim 1, wherein said membrane consists of a layer of biocompatible non-woven polymer located between two layers of porous biocompatible polymers.
 3. The encapsulating chamber according to claim 1, wherein said non-woven biocompatible polymer is chosen from the group consisting of polycarbonate (PC), polyethyleneimine, polypropylene (PP), poly(ethylene terephthalate) (PET), poly(vinyl chloride) (PVC), polyamide and polyethylene (PE).
 4. The encapsulating chamber according to claim 1, wherein said porous biocompatible polymer of at least one layer is chosen from the group consisting of polycarbonate (PC), polyethyleneimine, polypropylene (PP), poly(ethylene terephthalate) (PET), poly(vinyl chloride) (PVC), polyamide and polyethylene (PE).
 5. The encapsulating chamber according to claim 1, wherein at least one, or the two layer(s) of porous biocompatible polymer is (are) made hydrophilic by surface physical or chemical modification, and covered with at least one hydrophilic polymer.
 6. The encapsulating chamber according to claim 1, wherein one of the two layers of porous biocompatible polymers has a pore density of between 10⁶ pores/cm² and 10¹¹ pores/cm².
 7. The encapsulating chamber according to claim 1, wherein the total thickness of the membrane is between 45 μm and 200 μm.
 8. The encapsulating chamber according to claim 1, wherein the thickness of one of the layers of biocompatible polymer is between 5 and 40 μm, and the thickness of the other layer of biocompatible polymer is between 25 and 100 μm.
 9. The encapsulating chamber according to claim 1, wherein the internal diameter of the pores present on one of the layers of biocompatible polymer is between 5 and 100 nm, and the internal diameter of the pores present on the other layer of biocompatible polymer is between 100 and 2000 nm.
 10. The encapsulating chamber according to claim 5, wherein at least one layer is covered with a hydrophilic polymer which contains at least one biologically active molecule.
 11. The encapsulating chamber according to claim 1, further comprising a biocompatible sheet contained in said shell, said sheet optionally comprising projections at its surface.
 12. The encapsulating chamber according to claim 1, wherein the layer external to the shell has pores with an internal diameter of between 100 and 2000 nm, and the layer internal to the shell has pores with an internal diameter of between 5 and 100 nm.
 13. The encapsulating chamber according to claim 1, which comprises at least one connector which makes it possible to establish a communication between the exterior and the interior of the shell.
 14. The encapsulating chamber according to claim 1, which is circular and has a diameter of between 3 cm and 20 cm.
 15. A bioartificial organ, comprising at least one encapsulating chamber according to claim 1, in which secreting cells producing at least one substance of therapeutic interest are present.
 16. The bioartificial organ according to claim 15, which is a bioartificial pancreas containing insulin-secreting cells or islets of Langerhans.
 17. A process for obtaining an encapsulating chamber according to claim 1, comprising a step of heat-welding one or two membranes comprising a layer of biocompatible non-woven polymer located between two layers of porous biocompatible polymers so as to form a closed pouch intended to receive the secreting cells producing at least one substance of therapeutic interest.
 18. The encapsulating chamber according to claim 1, wherein said non-woven polymer is polyester.
 19. The encapsulating chamber according to claim 1, wherein said porous biocompatible polymer of at least one layer is polyester.
 20. The encapsulating chamber according to claim 1, wherein the internal diameter of the pores present on one of the layers of biocompatible polymer is between 5 and 100 nm, and the internal diameter of the pores present on the other layer of biocompatible polymer is between 200 and 1000 nm.
 21. The encapsulating chamber according to claim 2, wherein said non-woven biocompatible polymer is chosen from the group consisting of polycarbonate (PC), polyethyleneimine, polypropylene (PP), poly(ethylene terephthalate) (PET), poly(vinyl chloride) (PVC), polyamide and polyethylene (PE).
 22. The encapsulating chamber according to claim 2, wherein said porous biocompatible polymer of at least one layer is chosen from the group consisting of polycarbonate (PC), polyethyleneimine, polypropylene (PP), poly(ethylene terephthalate) (PET), poly(vinyl chloride) (PVC), polyamide and polyethylene (PE).
 23. The encapsulating chamber according to claim 2, wherein at least one, or the two, layer(s) of porous biocompatible polymer is (are) made hydrophilic by surface physical or chemical modification, and covered with at least one hydrophilic polymer.
 24. The encapsulating chamber according to claim 2, wherein one of the two layers of porous biocompatible polymers, has a pore density of between 10⁶ pores/cm² and 10¹¹ pores/cm².
 25. The encapsulating chamber according to claim 2, wherein the total thickness of the membrane is between 45 μm and 200 μm.
 26. The encapsulating chamber according to claim 2, wherein the thickness of one of the layers of biocompatible polymer is between 5 and 40 μm, and the thickness of the other layer of biocompatible polymer is between 25 and 100 μm.
 27. The encapsulating chamber according to claim 2, wherein the internal diameter of the pores present on one of the layers of biocompatible polymer is between 5 and 100 nm, and the internal diameter of the pores present on the other layer of biocompatible polymer is between 100 and 2000 nm.
 28. The encapsulating chamber according to claim 23, wherein at least one layer is covered with a hydrophilic polymer which contains at least one biologically active molecule.
 29. The encapsulating chamber according to claim 2, further comprising a biocompatible sheet contained in said shell, said sheet optionally comprising projections at its surface.
 30. The encapsulating chamber according to claim 2, wherein the layer external to the shell has pores with an internal diameter of between 100 and 2000 nm, and the layer internal to the shell has pores with an internal diameter of between 5 and 100 nm.
 31. The encapsulating chamber according to claim 2, which comprises at least one connector which makes it possible to establish a communication between the exterior and the interior of the shell.
 32. The encapsulating chamber according to claim 2, which is circular and has a diameter of between 3 cm and 20 cm.
 33. A bioartificial organ comprising at least one encapsulating chamber according to claim 2, in which secreting cells producing at least one substance of therapeutic interest are present.
 34. The bioartificial organ according to claim 33, which is a bioartificial pancreas containing insulin-secreting cells or islets of Langerhans.
 35. A process for obtaining an encapsulating chamber according to claim 2, comprising a step of heat-welding one or two membranes comprising a layer of biocompatible non-woven polymer located between two layers of porous biocompatible polymers, so as to form a closed pouch capable of receiving the secreting cells producing at least one substance of therapeutic interest.
 36. The encapsulating chamber according to claim 2, wherein said non-woven polymer is polyester.
 37. The encapsulating chamber according to claim 2, wherein at least one of the layers of porous biocompatible polymer of is polyester.
 38. The encapsulating chamber according to claim 2, in which the internal diameter of the pores present on one of the layers of biocompatible polymer is between 5 and 100 nm, and the internal diameter of the pores present on the other layer of biocompatible polymer is between 200 and 1000 nm. 